When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the excited nuclei in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) that is in the x-y plane and that is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mt. A signal is emitted by the excited nuclei or “spins”, after the excitation signal B1 is terminated, and this signal may be sampled and processed to form an image.
When utilizing these “MR (magnetic resonance)” signals to produce images, magnetic field gradients (Gx, Gy and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received MR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
The measurement cycle used to acquire each MR signal is performed under the direction of a pulse sequence produced by a pulse sequencer. Clinically available MRI systems store a library of such pulse sequences that can be prescribed to meet the needs of many different clinical applications. Research MRI systems include a library of clinically proven pulse sequences and they also enable the development of new pulse sequences.
The MR signals acquired with an MRI system are signal samples of the subject of the examination in Fourier space, or what is often referred to in the art as “k-space”. Each MR measurement cycle, or pulse sequence, typically samples a portion of k-space along a sampling trajectory characteristic of that pulse sequence. Most pulse sequences sample k-space in a roster scan-like pattern sometimes referred to as a “spin-warp”, a “Fourier”, a “rectilinear” or a “Cartesian” scan. The spin-warp scan technique is discussed in ‘Edelstein W A, Hutchison J M S, Johnson G, Redpath T. K-space substitution: Spin-Warp MR Imaging and Applications to Human Whole-Body Imaging. Physics in Medicine and Biology 1980; 25:751-756’. The method, referred to as spin-warp imaging, employs a variable amplitude phase encoding magnetic field gradient pulse prior to the acquisition of MR spin-echo signals to phase encode spatial information in the direction of this gradient. In a two-dimensional implementation (2DFT), for example, spatial information is encoded in one direction by applying a phase encoding gradient (Gy) along that direction, and then a spin-echo signal is acquired in the presence of a readout magnetic field gradient (Gx) in a direction orthogonal to the phase encoding direction. The readout gradient present during the spin-echo acquisition encodes spatial information in the orthogonal direction. In a typical 2DFT pulse sequence, the magnitude of the phase encoding gradient pulse Gy is incremented (ΔGy) in the sequence of measurement cycles, or “views” that are acquired during the scan to produce a set of k-space MR data from which an entire image can be reconstructed.
The role of the polarizing field B0 in the entire MR imaging process is indispensable as it influences the intrinsic signal to noise ratio (SNR) in the processed MR images. In an ideal scenario, the B0 field is identical at all spatial location within an imaging volume. A uniform B0 field optimizes SNR throughout the imaging volume, enables high resolution imaging, and facilitates specialized applications such as Magnetic Resonance Spectroscopy. However, while providing a uniform B0 field, the magnet sub-system of an MRI scanner introduces its own set of trade-offs into the overall MRI scanner design. Firstly, a uniform B0 field requires a bulky and expensive MRI magnet that typically requires a specialized room for its installation. Secondly, the structure of the magnet cannot be altered in any substantial manner due to the restriction placed on it by the uniformity specification of the B0 field. If the uniformity of B0 is degraded to address these concerns, the magnetic field gradient amplitudes have to be increased to preserve high resolution imaging capabilities.
MRI systems and methods for imaging in inhomogeneous fields have been proposed before. For example, in U.S. Pat. No. 5,304,930, a device entitled “Remotely Positioned MRI System” was described, which is an example of an MRI device which has a non-homogeneous static magnetic field. Another example of an MRI system designed to generate a non-homogeneous magnetic field has been described in EP 0 887 655 B1. In both these examples, the objective was to alter the design of the magnet itself to improve access to the patient during imaging. In contrast, in high field MRI scanner with homogeneous fields, the size of the opening for patient access to the magnetic field in these devices is generally restricted in size. This limited access results from the fact that producing a high degree of field homogeneity places inherent restrictions on the structural dimension of an MRI magnet.
The approach in the two cited references was more focused on altering the structure of the magnet itself to obtain great flexibility in design at the expense of homogeneity. This approach while addressing the restrictive effects of the magnet structure cannot compensate for the resultant field inhomogeneity, which degrades image quality. Alternatively, reconstruction algorithms [1] have been proposed to compensate for the effects magnetic field inhomogeneity in what are otherwise homogeneous fields set up by cylindrical magnets. But these algorithms are designed to primarily address the issue of susceptibility induced artifacts that cause distortions in air/tissue interfaces. These algorithms cannot facilitate any changes in the restrictive structure of the MRI magnet itself.
Another approach [20] discloses techniques to correct geometric distortion seen in MR images obtained from employing pulse sequences that are sensitive to even mild (less than 5 ppm) magnetic field inhomogeneity that are generated by current, state of the art MRI magnets. The correction techniques discussed therein require that even this mild level of inhomogeneity be accurately modeled.
Therefore, the present invention seeks to address the above shortcomings experienced for MR imaging, especially for those captured in a highly inhomogeneous field. The present invention is capable of compensating for both mild (such as those discussed in [20]) and significant (greater than 10 ppm) levels of field inhomogeneity regardless of the type of pulse sequence that is employed.